The present invention relates, generally, to systems and methods for magnetic resonance imaging (MRI) and, more particularly, to systems and methods for measuring and quantifying substances in a subject using magnetic susceptibility.
Iron is an essential nutrient for the human body, needed by every human cell. However, excessive iron is toxic, and the body has very limited capabilities to eliminate abnormal accumulation of iron. There are a variety of conditions that afflict patients that require regular blood transfusions. For example, hese patients are at risk of developing “transfusional” iron overload. Iron overload can result in multiple complications, such as preventing normal growth and sexual maturation, and damaging the liver and heart. Treatment for transfusional iron overload is based on administering iron-reducing “chelator” agents (chemicals that bind to excess iron and remove it from the body), either orally or intravenously. Treatment with chelators is extremely expensive (>$40,000/year) and is potentially toxic. Therefore, accurate measurement of body iron levels is critical to determine when to initiate treatment. It is also important to monitor treatment (allowing the adjustment of chelator dose to maintain low iron levels while minimizing risks from the treatment).
The simplest method available to assess body iron is based on measuring serum ferritin or serum transferring receptor concentration from a blood sample. Unfortunately, it is well known that serum biochemical tests are often confounded by a number of factors and do not accurately reflect body iron levels. Hepatic iron content (HIC) is widely considered the best reference for assessing total body iron stores, because the amount of iron in the liver is closely correlated to total body iron. Currently, chemical analysis of liver biopsy samples is the best available reference standard to measure HIC. However, biopsy is limited because it is invasive, expensive ($1500-2000), and has poor sampling variability. In addition, biopsy cannot be performed in patients who have low platelets or low coagulation factors (e.g., myelodysplastic syndrome), due to the risk of uncontrolled bleeding.
Liver iron susceptometry using a superconducting quantum interference device (SQUID) is generally regarded as the most accurate non-invasive method to quantify liver iron. SQUID is a well-validated non-invasive reference standard for measuring HIC. Importantly, SQUID directly measures tissue susceptibility, which is a fundamental property of all substances including tissue. Further, the relationship between susceptibility and tissue iron concentration is well understood. This relationship is well understood because iron is the only naturally occurring non-trace substance in the body with significant magnetic susceptibility, and increases in local tissue susceptibility only occur from tissue iron overload. This fact, and direct relationship between magnetic susceptibility and iron concentration makes this a highly attractive and fundamental quantitative biomarker of tissue iron overload. Unfortunately, even though SQUID has been calibrated, validated and used for clinical studies, its complexity, high cost and limited availability (only four SQUID devices are available world-wide for this purpose) have precluded its widespread use.
Magnetic Resonance Imaging (MRI) is a widely available and accessible technology that has been shown to be very sensitive to the presence of iron. When a substance, such as human tissue, is subjected to a sufficiently-large, uniform magnetic field (polarizing field B0), the individual magnetic moments of the nuclei in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B1) that is in the x-y plane and that is near the Larmor frequency, the net aligned moment, Mz, may be rotated, or “tipped”, into the x-y plane to produce a net transverse magnetic moment Mxy. A signal is emitted by the excited nuclei or “spins”, after the excitation signal B1 is terminated, and this signal may be received and processed to form an image.
When utilizing these “MR” signals to produce images, magnetic field gradients (Gx, Gy, and Gz) are employed. Typically, the region to be imaged is scanned by a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The resulting set of received MR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques.
To do so, the signals are often weighted in different ways to give preference to or consider different sub-signals or so-called contrast mechanisms. Two basic “contrast mechanisms” commonly utilized in MR imaging are the spin-lattice (or longitudinal or T1) relaxation time or spin-spin (or transverse or T2) relaxation time. However, there are a variety of other mechanisms for eliciting contrast in MRI, including R2*. Specifically, T2* is a quantity related to T2, but including dephasing effects. That is, T2* is a quantity related to spin-spin relaxation and, in addition, relating magnetic field inhomogeneities and susceptibility effects. Often, instead of T2*, these quantities are preferably expressed in terms of relaxation, or the inverse of the T2* time constant, represented as R2*.
Iron introduces microscopic inhomogeneities in the magnetic field, resulting in an increased rate of signal decay, measured by two MRI relaxation constants R2 and R2*. MRI-based iron quantification methods are based on either R2 or R2* relaxivity, using spin-echo or gradient-echo sequences, respectively.
An R2 method (commercially termed “Ferriscan”) has been shown to provide accurate measurements of HIC and is FDA approved. This approach is based on a monotonic, curvilinear empirical calibration curve linking R2 and HIC. However, Ferriscan has several important limitations that have severely curtailed its widespread use: 1) it is expensive ($300/patient, in addition to all other MRI related costs), 2) it is very slow (10-20 minutes of scanning in order to obtain adequate R2 measurements), and 3) may have limited dynamic range due to a nonlinear relationship of R2 with HIC. Additionally, Ferriscan results cannot be obtained immediately after the MRI scan, but require uploading the acquired images to Resonance Health (located in Claremont, Australia), and the HIC estimate is received within two business days.
R2*-based methods have the potential to overcome these limitations. Due to the high speed of gradient-echo pulse sequences, single breath-hold whole-liver R2* mapping is possible. In recent years, a linear relationship between R2* and HIC has been demonstrated. Further, it has been demonstrated that R2*-based methods can be confounded by several factors including the presence of fat, macroscopic susceptibility and magnitude based signal estimation. The presence of fat has been successfully addressed through the use of R2* mapping combined with chemical shift encoded water-fat separation methods that use multipeak spectral modeling of fat. When complex fitting is used, the bias in R2* estimates caused by non-zero noise at low signal levels are naturally avoided. Also, the use of weighted least squares fitting or non-linear least squares fitting performed part of the chemical shift encoded decomposition can be used to avoid bias at high R2* values. Finally, the augmentation in the signal decay and overestimation in R2* values that results from macroscopic magnetic field inhomogeneities has also be addressed. Correction for this effect can be performed by measuring the local gradient of the local field inhomogeneity map (“field map”). The field map is estimated as part of complex chemical shift encoded water-fat separation methods. Hernando D, Vigen K K, Shimakawa A, Reeder S B. R2* mapping in the presence of macroscopic B0 field variations. Magn Reson Med. 2012 September; 68(3):830-40. recently demonstrated that the additional signal decay that results from gradients in the field map can be removed, avoiding overestimation of R2*. Finally, there is a known dependence on magnetic field strength on R2 and R2*, which means that all R2 and R2* methods require calibration for different magnetic field strengths.
Despite the potential for both R2 and R2* based methods to quantify tissue iron concentration and the demonstration of a monotonic relationships between these parameters and tissue iron concentration, they both share a fundamental limitation. Specifically, the relationship between both R2 and R2*, and HIC is not well understood. For this reason, the use of both R2 and R2* to measure iron concentration requires measurement of empirical calibration curves. While this approach is often practical in many circumstances, it is fundamentally limiting, since the development of new methods may require repeated recalibration, which is expensive and often prohibitive.
Therefore, it would be desirable to have a system and method for measuring or quantifying iron or other substances in a subject that is accurate and repeatable without the need for special calibrations, such as a calibration curve specific to each system or the like.